Spectroscopy is a method for obtaining information on a molecular scale by the use of light. This information can be related to the rotational, vibrational and/or electronic states of the molecules probed as well as dissociation energy and more. The rotational and/or vibrational spectrum of a given molecule is specific for that molecule. As a consequence, molecular spectra in particular rotation and/or vibrational spectra are often referred to as ‘fingerprints’ related to a specific molecule. Information related to rotational, vibrational and/or electronic states of molecules can therefore be used to analyze a sample comprising a number of unknown molecular components, thereby obtaining knowledge about the molecular components in the sample.
The basis for a spectroscopic setup is a light source, e.g. a laser, which is used for illuminating a sample. The light from the light source (the incoming light) will interact with the sample, and often result in an alternation of the light which is transmitted through, emitted by, reflected by and/or scattered by the sample. By collecting the altered light and analyzing its spectral distribution, information about the interaction between the incoming light and the molecular sample can be obtained; hence information about the molecular components can be obtained.
The spectral distribution is typically measured by using a spectrophotometer. A spectrophotometer is an optical apparatus that works by separating the light beam directed into the optical apparatus into different frequency components and subsequently measuring the intensity of these components by using e.g. a CCD detector, a CCD array, photodiode or such.
The altered light reflecting interactions between the incoming light and the molecular sample can roughly be characterized as either emission or scattering if the light is collected in generally the reverse path from that on which it entered the sample. The emission signals have relatively broad spectral profiles as compared to scattering light signals, which normally display quite narrow spectral lines. One process often dominates over the other, but both processes can and most often will occur simultaneously. The intensity of the emitted light vs. the intensity of the scattered light depends among other things on the frequency and the power of the incoming light, the intensity of the incoming light at the measuring point in the sample, and the molecular components in the sample.
Scattering of light can be classified as being either elastic or inelastic and these are characterized by being spectroscopically very narrow signals. Elastic scattering is referred to as Rayleigh scattering, in which there is no frequency shift. Rayleigh scattering thus has the same frequency as that of the incoming light.
The most commonly known example of inelastic scattering is Raman scattering, in which there is an energy interchange between the molecule and the photons of the incoming light. The frequencies, i.e. the spectral distribution of the Raman scattered light will be different from that of the incoming light and uniquely reflect the specific vibrational levels of the molecule; hence it is a fingerprint spectrum. This can be used for identification of the molecular composition of the substance probed and/or the concentration of the specific molecules in the substance.
Raman scattering is a relatively weak process compared to e.g. Rayleigh scattering and fluorescence. Reduction of contributions from these other processes is thus desirable when collecting Raman scattered light. In addition, the intensity of the Raman scattered light depends strongly on the frequency and the intensity of the incoming light. If these are variable, it may therefore be essential to monitor power fluctuations in the incoming light if one is to receive reliable information about the distribution of molecular components in different samples and/or sample spots based on analysis of the collected Raman scattered light, depending on the precision needed. The same is true if the analysis of the molecular components in a sample and/or different sample spots is based on emission spectra.
Skin comprises a number of layers having different characteristics and containing different kinds of cells and structures. Various proposals for using Raman spectroscopy to measure glucose or other components in skin have been made, but none of these has to date provided a system which ensures that in any given individual collected light signals originate from below the stratum corneum by a margin that provides for a good measurement.
The skin surface is formed by the stratum corneum which consists mainly of cornified, dead, flattened skin cells and varies in thickness between individuals and between areas of the body. The concentration of components such as glucose in the interior of the stratum corneum is not in equilibrium with the interstitial fluid below the stratum corneum.
Often, it is desired to make transdermal measurements on the fingertip because of the ease with which the fingertip can be placed into the light path of a suitable instrument. However, individual variations in the thickness of the stratum corneum are relatively large in this area. Thus, the stratum corneum generally has a thickness of from 10-15 μm on most areas of the body but may be more than 10 times thicker on the palms and the soles. Finger print patterns also provide variation in the thickness of the stratum corneum over the surface of the fingertips.
WO2011/083111 describes a Raman spectrometer based apparatus for transdermal measurement of glucose which is set to derive Raman signals from a depth of 60 to 400 μm below the skin surface, typically by focusing the incoming light to a depth within the range of 200-300 μm. This is found to be broadly satisfactory, but fails with occasional subjects whose stratum corneum at the measurement site is too great.
Caspers et al; Biophysical Journal, Vol 85, July 2003, describes an in vivo confocal Raman spectroscopy method and apparatus which is said to be useful for measuring glucose. It contains however no instruction as to the depth from which the Raman scattering should be collected in a glucose measurement and there is a strong suggestion deducible from the teaching that the apparatus had not actually been tried for this purpose.
WO2008/052221 describes a method and apparatus for coherent Raman spectroscopy that transmits light through a sample surface such as skin and tissue to a focal plane within the sample to measure for instance glucose. However, no teaching is present of the importance of selecting a particular depth for the focal plane or where this should be. Indeed, it is specifically acknowledged that using the described apparatus variations in the detected signal occur when the analyte concentration is constant due to effects of skin temperature and hydration. No suggestion is present that such effects can be avoided by a careful selection of the depth from which the measurements are taken.
WO97/36540 describes determination of the concentration of e.g. glucose using Raman spectroscopy and an artificial neural network discriminator. However, the Raman signals are not selectively obtained from a particular depth and the need to compensate for non-linearities arising from signals penetrating to a depth of >500 μm is discussed.
WO00/02479 discloses a method and apparatus for non-invasive glucose measurement by confocal Raman spectroscopy of the aqueous humor of the anterior chamber of the eye. Naturally, there is no teaching of a depth at which to make optimal measurements in skin.
WO2009/149266 refers back to Ermakov I V, Ermakova M R, McClane R W, Gellermann W. Opt Lett. 2001 Aug. 1; 26(15):1179-81, ‘Resonance Raman detection of carotenoid antioxidants in living human tissues.’ which describes using resonance Raman scattering as a novel non-invasive optical technology to measure carotenoid antioxidants in living human tissues of healthy volunteers. By use of blue-green laser excitation, clearly distinguishable carotenoid Raman spectra superimposed on a fluorescence background are said to be obtained.
Chaiken et al (Noninvasive blood analysis by tissue modulated NIR Raman spectroscopy, J. Chaiken et. al., Proc. of SPIE optical Eng., 2001, vol. 4368, p. 134-145) obtained a correlation of only 0.63 between Raman based measurements and finger stick blood glucose measurements across several individuals, but were able to obtain a correlation of 0.90 for a single individual. The setup utilized by Chaiken et al comprises a collimated excitation beam and so naturally they do not disclose any optimal focal depth.
The present invention now provides a method for determining whether the origin of Raman signals received in transdermally operating confocal detector apparatus lies within the stratum corneum or below it, which method comprises analysing said signals to compare the relative intensities of Raman signals originating from a first skin component and Raman signals originating from a second skin component, wherein said relative intensities are indicative of whether the Raman signals originate within the stratum corneum or below the stratum corneum.
Alternatively expressed, the invention provides a method for predicting whether a spectrum of Raman signals received transdermally in a confocal detector apparatus and having at least one component expected to have an intensity representing the concentration of a skin component at a point of origin of said Raman signals below the surface of the skin will accurately represent said concentration, which method comprises analysing features of said spectrum relating to skin components other than the skin component the concentration of which is to be measured and thereby determining whether the Raman signals originate primarily within the stratum corneum so that the spectrum will be less likely to represent said concentration accurately or originate primarily below the stratum corneum so that the spectrum will be more likely to represent said concentration accurately.
The preferred features described below apply to either of these aspects of the invention.
Preferably, the method comprises analysing said signals to compare the relative intensities of Raman signals originating from a first skin component and Raman signals originating from a second skin component, wherein said relative intensities are indicative of whether the Raman signals originate primarily within the stratum corneum or primarily below the stratum corneum.
Preferably, said first skin component produces a peak in the Raman spectrum at a wavenumber of 883-884 cm−1. This may derive from proteins, including Type I collagen (see Raman Spectroscopy of Biological Tissues, Movasaghi et al, Applied Spectroscopy Reviews 42: 493-541, 2007).
Preferably said second skin component produces a peak in the Raman spectrum at 893-896 cm−1. This may derive from deoxyribose phosphodiester. Thus, the second skin component may be DNA.
The methods of the invention may further comprise the step of comparing the sizes of said first and second peaks and producing an output indicative that the signals arise from within the stratum corneum if the size of said first peak divided by the size of said second peak is less than a selected value R. The value of R may be chosen according to the selectivity of the determination of origin of the analyte signal that one wishes to achieve. It will generally be convenient to use the peak height as a measure of peak size but one may adopt another size measure such as area.
If R is made larger, more candidate measurements are likely to be rejected, leading to an increased need to repeat such measurements at a different measurement site, or to adjust the focussing distance of the apparatus used, or to reject the patient from this form of measurement entirely.
If R is chosen to be smaller, fewer measurements will be ruled unreliable, but the chances of a measurement being accepted that does not in reality correlate well to actual analyte concentration may be increased.
Preferably, R is selected to be at least 0.75, more preferably at least 0.95, and optionally R is set at 1.0 or higher, e.g. up to 1.25
Supposing that R is set to be 1.0, if the said 883-884 cm−1 peak is higher than the 893-896 cm−1 peak, this is a good indicator that the signals arise from sufficiently below the stratum corneum for the measurements of an analyte skin component in interstitial fluid to be accurate. If on the other hand the height order is reversed and the 893-896 cm−1 peak is higher than the 883-884 cm−1 peak, this is an indicator that the signals may not arise from sufficiently below the stratum corneum for the measurements of an analyte skin component in interstitial fluid to be accurate. However, it is expected that the value of 1.0 builds in a safety margin and R could be set lower.
Accordingly, the method may include the steps of comparing the sizes of said first and second peaks and producing an output indicative that the signals arise from within the stratum corneum if said first peak is of lesser size than said second peak and/or producing an output indicative that the signals arise from below the stratum corneum if said second peak is of lesser size than said first peak.
Other peaks in the Raman spectrum may be chosen that provide a similar indication of the depth of origin of the signals.
The above methods provide a first line test, but even if the said size relationship of the 883-884 cm−1 peak and the 893-896 cm−1 peak is satisfactory, this may not in all cases provide sufficient assurance and a second line check may be desirable. To this end, one may investigate whether the size of a Raman peak deriving from a third and/or the size of a Raman peak deriving from a fourth skin component, or further skin components, is greater than a predetermined size. Again, height may be adopted as a convenient measure of size.
The predetermined size for this purpose for each Raman peak used may be x standard deviations above the mean size for the respective peak as measured using the same apparatus on multiple different measurement sites on the skin of a test population of one or preferably multiple test subjects. Suitably, x may be from 0.5 to 2, e.g. 0.75 to 1.5, but is preferably 1.
Thus for instance, if the comparison of the relative intensities of Raman signals originating from a first skin component and Raman signals originating from a second skin component is indicative that the Raman signals originate primarily below the stratum corneum, the methods of the invention may further comprise determining whether the size of a further peak in the spectrum associated with a skin component which may be prevalent in the stratum corneum is more than one standard deviation greater than a mean value for the size of that peak in a statistically valid sample of similar spectra, a positive determination indicating a probability that the Raman signals do not after all originate primarily below the stratum corneum.
To obtain the similar spectra, multiple measurement sites are suitably chosen to be sufficiently numerous to provide a statistically valid measurement of the standard deviation. Suitably, 100-300 test sites, preferably chosen on from 5-20 test subjects, for instance 10 sites on each of 20 individuals could be used. The test subjects should preferably be matched for ethnicity, age, and/or nature of occupation (such as manual worker or not) with each other and the subject of the analyte measurements.
Suitable peaks for use in this second line check would be peaks at 1445 cm−1 and at 1650 cm−1. The former may arise from various bending modes associated with CH2 and CH3 groups in collagen. The latter may derive from protein amide groups.
If the first line test is failed or if either of these chosen peaks is larger than the chosen cut off size, the validity of the analyte measurement is doubtful and an alternative measurement site should be chosen or the depth from which the Raman signals originate should be suitably altered, which will generally imply that it should be increased.
The method of the invention can be used to determine whether the transdermally operating confocal detector apparatus will successfully measure the concentration of a target skin component by measuring Raman signals originating below the stratum corneum. This determination may be used simply to exclude patients from measurement who are unsuitable, or to guide a choice of a different measurement site where the stratum corneum is not too thick, or to guide an adjustment of the transdermally operating confocal detector apparatus in order to cause it to measure Raman signals originating from below the stratum corneum.
Methods according to either aspect of the invention may therefore further include adjusting said transdermally operating confocal detector apparatus in response to a finding that the Raman signals originate from within the stratum corneum, said adjustment altering the depth of origin of said Raman signals such that a new depth is determined to be satisfactory and in particular such that the depth is no longer determined to be within the stratum corneum.
This may be done by adjustment of a distance from the surface of the skin of an objective lens from which light is emitted to the measurement site and received from the measurement site, so as to alter the depth from which Raman signals are received.
Alternatively or additionally, said transdermally operating confocal detector apparatus may comprise an objective lens having a focal length and said method of altering the depth of origin of the Raman signals may comprise altering the focal length of the objective lens by replacement of the objective lens or by adjustment of the objective lens.
To this end, said transdermally operating confocal detector apparatus may comprise a compound objective lens comprising at least a first element and a second element spaced from the first element, and said method of altering the depth of origin of the Raman signals may comprise altering the spacing of two or more elements to adjust the focal length of the compound objective lens, and this would include replacement of the lens with one in which the said spacing is different.
Optionally, said adjustments of the lens position or focal length may be carried out by altering the thickness of a piezoelectric spacer, either between the objective lens and the skin surface or between said lens elements by the alteration of a voltage applied thereto.
Optionally, said adjustments of the lens position or focal length may be carried out by altering the rotational position of an annular screw mounted collar carrying at least one element of said compound lens.
The invention also provides in a further aspect transdermally operating confocal detector apparatus for non-invasive in vivo measurement by Raman spectroscopy of the concentration of a skin component present in the skin of a subject, comprising a light source, optical components defining a light path from said light source to a measurement location, a spectrum analysis unit, optical components defining a return path for Raman scattered light from said measurement location to said spectrum analysis unit, wherein said spectrum analysis unit operates to determine whether the origin of Raman signals received therein lies within the stratum corneum or below it, by analysing features of Raman scattered light relating to skin components other than the skin component the concentration of which is to be measured and thereby determining whether the Raman signals originate primarily within the stratum corneum or primarily below the stratum corneum.
Alternatively expressed, in this aspect the invention provides in a further aspect transdermally operating confocal detector apparatus for non-invasive in vivo measurement by Raman spectroscopy of the concentration of a skin component present in the skin of a subject, comprising a light source, optical components defining a light path from said light source to a measurement location, a spectrum analysis unit, optical components defining a return path for Raman scattered light from said measurement location to said spectrum analysis unit, wherein said spectrum analysis unit operates to determine whether the origin of Raman signals received therein lies within the stratum corneum or below it, by analysing said signals to compare the relative intensities of Raman signals originating from a first skin component and Raman signals originating from a second skin component, wherein said relative intensities are indicative of whether the Raman signals originate within the stratum corneum or below the stratum corneum.
Preferably, said spectrum analysis unit determines the size of a peak in the Raman spectrum at 883-884 cm−1 produced by said first skin component.
Preferably, said spectrum analysis unit determines the size of a peak in the Raman spectrum at 893-896 cm−1 produced by said second skin component.
Said spectrum analysis unit may determine a ratio between the size of a first peak in the Raman spectrum at 883-884 cm−1 and the size of a second peak in the Raman spectrum at 893-896 cm−1. Height may be used as a suitable measure of peak size.
Preferably therefore, said spectrum analysis unit produces an output indicative that the signals arise from within the stratum corneum if the height of said first peak divided by the height of said second peak is less than a selected value R. R may be pre-set to be 0.75, more preferably 0.95 and still more preferably 1.0.
Preferably, R is not more than 1.25.
Optionally, if the signal analysis unit determines that comparison of the relative intensities of Raman signals originating from a first skin component and Raman signals originating from a second skin component is indicative that the Raman signals originate primarily below the stratum corneum, said signal analysis unit further determines whether the size of a further peak in the spectrum associated with a skin component prevalent in the stratum corneum is more than x standard deviations greater than a mean value for the size of that peak in a statistically valid sample of similar spectra, a positive determination indicating a probability that the Raman signals do not after all originate primarily below the stratum corneum. The value of x is discussed above.
Said transdermally operating confocal detector apparatus may comprise a set of interchangeable objective lenses of differing focal length or an objective lens having an adjustable focal length. To this end, said objective lens may be a compound objective lens comprising at least a first element and a second element spaced from the first element, and said lens is then adjustable by altering the spacing of two or more elements to adjust the focal length of the compound objective lens. The spacing adjustment may preferably be piezoelectric or screw operated as described above. Interchangeable lenses may differ from one another in spacings of such lens elements.
The apparatus may include means for computing a concentration of glucose or another analyte component in interstitial fluid or blood based on analysis of said Raman scattered light. The Raman spectrum may be analysed by application thereto of a trained statistical model which relates peak intensities to glucose or other analyte concentration. This may be performed using partial least squares regression (PLS) as described in more detail in the references acknowledged in M. A. Arnold; In Vivo Near-Infrared Spectroscopy of Rat Skin Tissue with Varying Blood Glucose Levels; Anal. Chem. 2006, 78, 215-223 therein and in A. M. K. Enejder et al; Raman Spectroscopy for Non-invasive Glucose Measurements; Jnl of Biomedical Optics, 10(3), 031114; May/June 2005. Other forms of multivariate calibration may be used including Principal Component Analysis (PCA) in a manner analogous to that described in for instance A. G. Ryder, G. M. Connor and T. J. Glynn; Quantitative Analysis of Cocaine in Solid Mixtures using Raman Spectroscopy and Chemometric Methods; Journal of Raman Spectroscopy, 31; 221-227 (2000) or in J. T. Olesberg, L. Liu, V. V. Zee, and M. A. Arnold; In Vivo Near-Infrared Spectroscopy of Rat Skin Tissue with Varying Blood Glucose Levels; Anal. Chem. 2006, 78, 215-223. In general, statistical methods of spectrum analysis useful in calibrating detection of analytes from absorption spectra will be useful in analysis of Raman spectra also.
The apparatus may be adjustable to alter the depth below the skin surface from which most of the intensity of the Raman signals originates, so as to set said depth to be below the stratum corneum. Light collection will be from a range of depths and the apparatus may be adjustable such that a desired percentage of the light originates from, below the stratum corneum.
Preferably, said percentage is at least 55%, more preferably at least 70%, more preferably at least 90%. Preferably also, at least 90% of Raman scattered light received at the light detection unit originates at depths less than 600 μm below the skin surface. Preferably also less than 25%, more preferably less than 10%, of Raman scattered light received at the spectrum analysis unit originates at depths less than 100 μm below the surface of the skin.
Preferably, at least 40%, more preferably at least 50% of the light reaching the spectrum analysis unit originates from 200 to 400 μm below the surface of the skin.
Optionally, said adjustment is automated. Thus, the spectrum analysis unit may operate to determine the origin of the Raman signals as described and in the event of a determination that the signals originate from within the stratum corneum may output a control signal to an adjustment means which operates to adjust a confocal depth position from which Raman signals are received until the spectrum analysis unit determines the origin of the signals to be below the stratum corneum. Such adjustment means may produce a said control signal in the form of a voltage applied to a piezoelectric actuator for changing the position of at least one lens element relative to the skin in use. Alternatively, the control signal may drive a motor to rotate a rotatable lens adjustment mechanism to alter lens component spacing.
Thus the adjustment means may hunt for a satisfactory confocal depth by making iterative increases in the confocal depth until a satisfactory result is obtained. Suitably, this might entail making progressive increases in confocal depth of from 10 to 50 μm, e.g. from 20 to 30 μm.
The spectrum analysis unit may receive light from the surface of the skin without transmission of said light through an optical fibre or with such a fibre. In the latter eventuality, apparatus according to the invention may comprise a hand piece for application to the skin containing components defining said measurement location in use, and one or more optical fibres connecting said hand piece to said light source and to the spectrum analysis unit for analysis of signals received from said light detection unit to provide said measurement therefrom.
The position distal of a skin engaging member of said measurement location is optionally adjustable and can for instance be adjustable to be from 60 to 400 μm beyond said distal surface of the skin engaging member or can be adjusted to be from 50 to 400 μm, more preferably 200 to 300 μm, beneath the surface of the skin. Alternatively, however the position distal of the skin engaging member of said measurement location is fixed, suitably such that the numerical parameters discussed above are achieved.
Thus, the depth of focus of the optical components defining said light path, and/or the optical components defining said return path may be fixed rather than adjustable. It follows that in this case, if the spectrum analysis unit determines that the confocal depth is not satisfactory, an alternative measurement site should be chosen or that patient should be excluded.
The invention includes a method for non-invasive in vivo measurement by Raman spectroscopy of a component, which may be glucose, present in interstitial fluid in the skin of a subject, comprising, in either order, (a) directing light from a light source into the skin of said subject via optical components defining a light path from said light source to a measurement location in the skin and so producing Raman signals returning from the skin, determining whether the origin of the returning Raman signals lies within the stratum corneum or below it by analysing said signals to compare the relative intensities of Raman signals originating from a first skin component and Raman signals originating from a second skin component, wherein said relative intensities are indicative of whether the Raman signals originate within the stratum corneum or below the stratum corneum, and (b) directing light from said light source into the skin of said subject via said optical components defining a light path from said light source to said measurement location in the skin, receiving Raman scattered light back from the skin at a light detection unit via optical components defining a return path for Raman scattered light from said measurement location to said light detection unit, and determining said concentration from said Raman scattered light. Preferred features of step (a) above may be as previously described.
The method may further include adjusting said optical components so that the Raman signals are determined to originate from below the stratum corneum.
Such methods are preferably performed using apparatus in accordance with the invention.
The method may include calibrating the output of the apparatus by the use of the apparatus to provide an output in respect of a known concentration of the skin component to be measured prior to said measurement on said subject. Once calibrated the apparatus preferably is not calibrated again for a period of not less than a week, more preferably a month. Preferably, said calibration step of providing an output in respect of a known substance concentration is not carried out by the use of the apparatus on said subject.
Thus, the calibration may be conducted on a different subject for whom a concentration of the component is known or may be conducted using a standard reference material such as a drop of component solution placed in the measurement location or a solid phantom simulating a component solution.
Any apparatus described herein may be used in such a method.
The light source is preferably a laser. A preferred form of laser to use as the light source is a diode laser with a wavelength in the range of 300-1500 nm. Suitable preferred wavelengths are 785, 830, or 850 nm. A suitable power range is 50-1000 mW. For example, one may use an 830 nm, 500 mW FC-830 laser from RGB Lase.
The apparatus may include an optical probe for measuring light signals in which the optical components defining the light path from the light source to the measurement location comprise a first optical fibre guiding incoming light from said light source and a lens focusing said incoming light towards, i.e. into or onto, the measurement location. The optical components for defining a return path for Raman scattered light may comprise said lens and further optical components guiding the altered light to the spectral analysis unit. The further optical components may include a second optical fibre, however, instead of employing a second optical fibre, a spectrophotometer may be integrated directly into the handpiece. Optionally, there may be a further light detection unit (or light logging device) measuring intensity fluctuations in said incoming light, and this further light detection unit may advantageously be positioned after said first optical fibre, whereby said further light detection unit receives a part of said incoming light from said first fibre.
An electrical output from this light logging device which is representative of the intensity of the incoming light may be used to adjust intensity measurements in the spectrum analysis unit to compensate for variations in said intensity.
The use of at least one optical fibre is advantageous in that although a microscope can be used, a microscope-based optical probe is not a readily movable object and a user's body part could be awkward to place in a position where measurements could be made. A possibility would be for the patient to insert his/her finger or arm directly under or above the microscope objective in the microscope. Unfortunately, this is cumbersome if not impossible with most microscopes.
An optical probe employing not the whole microscope but only microscope objective(s) mounted separately on e.g. a table allows for better accessibility between probe and sample. Measurements of blood sugar levels or other skin components in a patient in vivo become more convenient as the patients arm or finger can be placed in front of the microscope objective(s) without much difficulty. However, if the chosen sample is a leg, it might prove more difficult to place it appropriately in front of the microscope objective(s).
Inside the optical probe, said light logging device will normally be positioned after a dichroic mirror, which allows a minor part of the incoming light to either pass through the dichroic mirror and onto said light logging device or to be reflected by the dichroic mirror onto said light logging device. Alternatively, a splitting device can be positioned between said first fibre and said dichroic mirror, where said splitting device reflects a minor part of the incoming light onto said light logging device.
One advantage with using a light logging device is that it allows for a precise measure of the variations in the intensity of the incoming light at all material times. This ensures that variations in the intensity of the altered light due to variations in the incoming light and not sample variations can be compensated for.
In an embodiment of the invention, said lens focusing incoming light towards said sample is arranged at the surface of said optical probe such that said lens is in direct contact with the skin during measuring.
An advantage with having the lens in direct contact with the skin during measurement is that the sample penetration depth, and thereby the distance from the optical probe to the sample focus point, is known exactly, as it is defined by the focal length of the lens.
In another embodiment of the invention, said optical probe further comprises a window, where said window is positioned between said lens and the skin, such that said window is in direct contact with the skin during measuring, and where the thickness of said window is smaller than the focal length of said lens.
An advantage with inserting a window between the lens and the skin is that it can provide an easier cleaning of the optical probe, if a fragile lens sensitive to cleaning is used.
Another advantage with inserting a window between the lens and the skin is that the penetration depth can be varied depending on the thickness of the window. This provides one way of setting the penetration depth to a value resulting in a determination that the Raman signals measured originate below the stratum corneum.
Equally, instead of having a solid window, a window aperture can be provided between the lens and the skin, the aperture being formed in a skin engaging member.
The optical probe according to the invention, may further comprise a dichroic mirror positioned after said first optical fiber, where said dichroic mirror reflects any percent between re_in=0 and 100 (e.g. 90%) and transmits any percent between tr_in=0 and 100 (e.g. 10%) of said incoming light, where re_in+tr_in=100 percent (ignoring losses), and reflects any percent between re_se=0 and 100 (e.g. 30%) and transmits any percent between tr_se=0 and 100 (e.g. 70%) of said altered light, where re_se+tr_se=100 percent (ignoring losses). Hence said dichroic mirror may reflect most of the incoming light and transmit most of the altered light.
Said dichroic mirror is normally positioned at an angle of 45 degrees in relation to the propagating direction of said incoming light out of said first optical fibre.
In an embodiment where most of the incoming light is reflected by the dichroic mirror, said light logging device may be positioned after said dichroic mirror, whereby said light logging device measures intensity fluctuations in said incoming light transmitted through said dichroic mirror.
In another embodiment where most of the incoming light is reflected by the dichroic mirror, a splitting device may be positioned between said first optical fibre and said dichroic mirror, whereby said light logging device measures intensity fluctuations in said incoming light reflected of by said splitting device.
In an embodiment of the invention, said dichroic mirror is transmitting most (e.g. ≥90%) of the incoming light whilst passing a minor portion (e.g. ≤10%) and is reflecting most of the altered light (e.g. ≥70%) whilst passing a smaller amount (e.g. ≤30%).
In an embodiment where most of the incoming light is transmitted by the dichroic mirror, said light logging device may be positioned after said dichroic mirror, whereby said light logging device measures intensity fluctuations in said incoming light reflected by said dichroic mirror.
An advantage of having the light logging device situated directly after said dichroic mirror is that it utilizes the part of the incoming light, which is not reflected by the dichroic mirror, and otherwise would be lost. There is consequently no need for any additional optical components to be inserted inside the optical probe in order collect light for measuring of the fluctuations in the incoming light.
In one embodiment of the invention, the angle α between the direction of light out of said first optical fibre and the direction of light entering a said second optical fibre is substantially α=90 degrees. The angle could also be in the range α=80-100 degrees.
In one embodiment of the invention, said optical probe further comprises at least a first aperture where said first aperture only allows altered light from the focus point in the skin to reach the spectrum analysis unit thereby ensuring depth confocality. Said aperture can be a separate element, but a narrow opening of a second fiber can equally well function as said aperture when a second fibre is used.
An advantage with using an optical aperture positioned before the spectrum analysis unit is that the optical aperture works as a 3D depth filter eliminating optical signals generated outside of the confocal depth, i.e. the sample focus spot. The advantage with using a confocal optical probe is that the altered light entering the spectrum analysis unit arises solely from interactions between the incoming light and the skin at the focus depth; hence contributions from the cone-like areas above and below the focus spot are minimized or eliminated.
In another embodiment of the invention, one or more apertures can additionally be employed to obtain a sharper contrast in the z (depth) direction. A second aperture is preferably positioned between the skin and the lens focusing the light into the sample. This second aperture can be separate element, but a narrow opening of the optical probe at the point where light exits/is collected by the lens can equally well function as an aperture.
Although apparatus according to the invention is designed and configured for measuring optical signals in the skin in vivo, it could also be employed for measuring optical signals by immersing it into e.g. a blood sample thereby making the measurement in vitro.
Generally, the optical elements found inside an optical probe of apparatus according to the present invention are enclosed by a cover. A preferred optical probe can be moved around freely due to the use of a flexible fibre for guiding light into and optionally out of the optical probe. This enables easy in vivo measurements of e.g. blood sugar levels in a patient using different body areas such as an arm, a finger, a leg or similar. The apparatus may however be constructed so that the optical components are contained in a housing which defines a specific location on which to place a fingertip pad for performance of the measurement.
The stratum corneum thickness of a fingertip pad will typically be from 10-40 μm (see Marks, James G; Miller, Jeffery (2006). Lookingbill and Marks' Principles of Dermatology (4th ed.). Elsevier Inc. Page 7. ISBN 1-4160-3185-5 and Thickness of the Stratum Corneum of the Volar Fingertips H. FRUHSTORFER, U. ABEL, C.-D. GARTHE, AND A. KNU″TTEL. Accordingly, the preferred measurement depths of 200-300 μm will be from 160 to 190 μm up to 260 to 290 μm below the stratum corneum. Depths of measurement for all skin areas are preferably from 50 to 390 μm, more preferably from 190 to 290 μm below the stratum corneum.
A primary application of the apparatus is generally to measure blood sugar levels in a patient. The level of glucose in blood correlates with the level in interstitial fluid at the selected depth. Other analytes which may be measured in the same way would include lactate, haemoglobin, cholesterol, alcohol, urea and/or drug.